Various mechanical instruments have been devised for microsurgery, such as in eye surgery, where the removal of the gel-like vitreous and associated membranes is a difficult and delicate task in the treatment of vitreoretinal diseases. Other microsurgical procedures include the formation of a new drainage fistula in the sclera of glaucoma patients to relieve the intraocular pressure, removal of lens material in cataract surgery, removal of masses from herniated discs in back surgery, and the partial removal of soft tissues in joint and brain surgery. The inherent disadvantages of the mechanical devices are caused by the frictions of the movements necessary for cutting and the traction and shearing forces exerted upon adjacent tissues. Such mechanical disturbances of delicate structures often result in excessive scarring, which can reduce or even eliminate the benefit of the operation, or which can delay the healing process.
Recent advances in technology and techniques have made the use of lasers in surgical procedures increasingly common. The precision attained through the use of lasers and laser equipment has particular advantages in microsurgical procedures which were formerly either not possible, or more traumatic using conventional instruments. As the field has developed, however, some of the techniques which are now in use and which are theoretically sound have led to previously unforeseen problems.
The pulsed mid-infrared lasers, such as Thulium (1.96 um wavelength, 100 um penetration depth in water), Holmium (2.1 um, 300 um penetration depth), and Erbium:YAG (2.94 um, 1 micron penetration depth), as well as the ultraviolet excimer lasers from 193 nm to 308 nm have been shown to be capable of ablating biological tissues with minimal thermal damage to the remaining structures. The size of the thermal damage depends on the penetration depth of the specific wavelength and pulse duration: the shorter these parameters, the smaller the zone of thermal damage. Since ablation is a threshold effect, the low energy density at the wings of a tapered laser beam profile may not reach the amount necessary to cause ablation, but only heat this part of the irradiation site, contributing to thermal damage. A homogeneous cross-sectional energy distribution is therefore highly desirous in order to reduce the lateral size of the thermal damage, which is one of the stimuli for undesired scarring and delayed wound healing.
Reduction of thermal damage by limiting the laser pulse duration to the thermal relaxation time has been proposed before by Wolbarsht. The thermal relaxation time has been connected to the penetration depth of the specific laser wavelength by defining it as that time, during which the major portion of the laser energy which is absorbed in the penetration depth, diffuses by heat conduction into a heat diffusion zone--which is proportional to the zone of thermal damage--of the same size as said penetration depth. This has two major inconveniences. First, in the case of the Holmium laser, where the penetration depth (in water) is about 300 um, the allowable maximal laser pulse duration becomes 156 milliseconds, which would result in a relatively large lateral heat diffusion zone of 300 microns. Second, in the case of the Erbium:YAG laser with a penetration depth of 1 micron, it limits the maximal pulse duration to 1.7 microseconds. New considerations are necessary to select more practical pulse durations and to allow zones of heat diffusion which are better adapted to the specific microsurgical procedure.
The Erbium:YAG laser energy can be transmitted through zirconium fluoride optical fibers. These fibers are relatively fragile, and the fiber output end is often damaged during tissue ablation. Provisions have to be taken to avoid the contact of this fiber with biological tissue. The energy of the 193 nm excimer laser can not be transmitted through optical fibers, and articulated arms have to be used to direct the laser beam into a probe.
The small penetration depth of mid-infrared and ultraviolet lasers, which causes the high precision in tissue ablation, inherently limits the amount of ablated tissue and the efficiency of such procedures. It is therefore advantageous to increase the diameter of the irradiated area by using bundles of optical fibers. Laser energy delivery in an annular pattern with a coaxial canal has been proposed before by Berlin in U.S. Pat. No. 4,846,172 and by Davies in U.S. Pat. No. 4,819,632. These patents teach arranging multiple optical fibers in a ring-like pattern and surrounding a central tube. These systems have the disadvantage of needing a large number of small and densely packed fibers to approximate a regular annular pattern, which, in cases involving a short penetration depth of the wavelength, will always consist of multiple spots rather than a homogeneous annulus. In addition, the holding fixture for such a fiber bundle increases the outer diameter of the probe by an undesired amount beyond the diameter of the annulus formed by the laser energy.
Single pulses of mid-infrared lasers are capable of evaporating water and tissue depths in excess of the optical penetration depth, provided the pulse duration and power density are long and high enough to induce a phase change in said penetration depth in a time t.sub.v, which is shorter than the pulse duration t.sub.p. Later portions of such pulses will then be transmitted through the vapor, which was produced by the earlier portions of the pulse. In conclusion, the laser energy of all mid-infrared lasers can travel a considerable distance beyond the output surface of any probe which is submersed in water or brought in contact with the tissue.